Magnetic resonance (MR) imaging is a known technology that can produce images of the inside of an examination subject without radiation exposure. In a typical MR imaging procedure, the subject is positioned in a strong, static, homogeneous base magnetic field B0 (having a field strength that is typically between about 0.5 Tesla and 3 Tesla) in an MR apparatus, so that the subject's nuclear spins become oriented along the base magnetic field.
Radio-frequency (RF) excitation pulses are directed into the examination subject to excite nuclear magnetic resonances, and subsequent relaxation of the excited nuclear magnetic resonances can generate RF signals. Rapidly switched magnetic gradient fields can be superimposed on the base magnetic field, in various orientations, to provide spatial coding of the RF signal data (also referred to as image data). The RF signal data can be detected during a ‘readout’ phase, and mathematically processed to reconstruct images of the examination subject. For example, the acquired RF signal data are typically digitized and stored as complex numerical values in a k-space matrix. An associated MR image can be reconstructed from the k-space matrix populated with such values using a multi-dimensional Fourier transformation.
One use of magnetic resonance imaging is to visualize myocardial damage in the setting of heart disease. Inversion-recovery (IR) and phase-sensitive inversion-recovery (PSIR) techniques, described in more detail below, can be used to visualize myocardial infarction and scar tissue in the setting of ischemic and non-ischemic heart disease. These techniques are also referred to as myocardial delayed enhancement (MDE) or late Gadolinium enhancement (LGE) sequences, and are widely used in clinical MRI.
Inversion recovery (IR) imaging is an MR technique that can provide T1 contrast between different tissue types. A conventional IR pulse sequence is shown in FIG. 2. In the IR sequence, an IR pulse (that is typically spatially and chemically non-selective), at the time labeled IR in FIG. 2, inverts longitudinal magnetization Mz of different tissue types, e.g., from +M0 to −M0. This may be done in the presence or absence of a T1-shortening contrast agent. The magnetization recovery (e.g., relaxation of the inverted magnetization from −M0 towards +M0) is different in each tissue type. For example, the species of primary interest in cardiac MRI are typically normal myocardium, infarcted myocardium (infarct), and blood. These different tissue types are often referred to as ‘species.’ T1 shortening contrast agents can optionally be introduced into the subject prior to the imaging procedure to improve contrast. The recovery rate of a particular species can be represented as the inverse of its longitudinal recovery time T1. For example, in FIG. 2, it can be seen that the longitudinal magnetization Mz of infarct tissue relaxes back to +M0 in a shorter time (i.e., with a shorter T1 value) than normal myocardium tissue, which has a longer T1 value.
A time delay TI is inserted between the IR pulse and the data acquisition sequence, such that the magnetization of one of the species (e.g., normal myocardium in the specific case of cardiac MRI) is approximately zero at the time of acquisition. This is often referred to as ‘nulling’ the particular species, and can generate improved T1 image contrast between the nulled species and other tissue types. For example, in FIG. 2, the normal myocardium magnetization recovery curve (shown as a dashed curved line) goes through zero at time TI after the initial inversion pulse (e.g., at or near the center of the inversion data acquisition sequence).
The scanner operator can set the so-called inversion time (TI) for the species to be nulled, which represents a time interval between the IR pulse and a time within the readout sequence where the image-contrast-relevant raw data is collected as illustrated, e.g., in FIG. 2. The TI shown in FIG. 2 effectively nulls the normal myocardium tissue such that it has essentially zero longitudinal magnetization, whereas the infarct and blood magnetizations (each of which has a shorter T1) have recovered to a greater degree and have positive M0 values at time TI following the IR pulse. In the exemplary IR image of a portion of a heart shown in FIG. 3 (with a TI set to the optimal TI time shown in FIG. 2), normal myocardium appears dark and infarct tissue appears bright.
In FIG. 2, the normal myocardium magnetization recovery curve shown in dashed line goes through zero at approximately the center of the raw data acquisition following the inversion pulse. This conventional sequence for acquiring raw IR image data is also referred to herein as an “inversion recovery data acquisition” or IR-DA. The optimal/“correct” TI value nulls normal myocardium during the IR-DA and, due to the shorter T1 of the infarct and blood, those species recover faster towards M0 (the initial baseline magnetization) and have a positive value at time TI following the IR pulse.
IR MR images are typically magnitude images, meaning the image pixels do not have a sign associated with them, even though the detected magnetization has a sign. Magnetizations with opposite polarity (sign) but the same magnitude appear with the same brightness in the in the MR image. For example, if TI were set at the time with poor infarct imaging denoted in FIG. 2, the infarct would appear black (having zero magnetization) and the normal myocardium would appear grey, having a moderate negative magnetization. Such inverted or ‘wrong’ image contrast can occur if the time delay between IR and data readout is set too short. More importantly, if TI were set between the two time values shown in FIG. 2 (e.g. between the ‘poor’ infarct imaging and the optimal time), there would be very little contrast between normal myocardium and infarct in the resulting IR image. For example, if normal myocardium has a slightly negative magnetization and infarct has a slightly positive magnetization at the selected TI, both types of tissue would appear as similar grey regions in an IR image because an IR image does not differentiate between positive and negative magnetizations. Accordingly, use of an incorrect TI value can result in uninterpretable or misinterpreted clinical images.
In cardiovascular imaging, the image data for a single MR image can take some time to be collected. Because the heart is beating and undergoes significant shape changes as it beats, cardiovascular images are often obtained by ‘triggering’ (also referred to as ‘gating’) a data collection sequence in sync with the heartbeat. This is often done using an electrocardiogram (ECG) signal, which detects the electrical activity of the heart. The electrical pulse peak that triggers the heart to begin a heartbeat is referred to as an R-wave. The time between consecutive R-waves can be referred to as a cardiac cycle, having the duration of a single heartbeat, and may be abbreviated as RR.
The pulsatile flow of blood from the heart with each heartbeat can also affect the shape and location of tissues outside the heart. For example, the aorta changes noticeably in shape and location during the cardiac cycle. The systems and methods described herein are therefore not limited to imaging cardiac tissue, but can also be used for extracardiac tissues that are affected by the periodic pulsatile motion of blood.
A schematic illustration of a conventional triggered IR data acquisition sequence is shown in FIG. 4. The upper squiggly line represents the ECG signal, with the R-wave pulses denoted by R. The interval labeled RR between consecutive R-waves represents a single cardiac cycle. In such triggered data acquisitions, the IR pulse (an RF pulse that typically inverts the magnetization of all tissues in the region or volume being imaged) is provided at some point in the cardiac cycle. Following a time delay TI, the inversion recovery data acquisition sequence (IR-DA) is performed, as also shown in FIG. 2. Additional IR data generation/acquisition sequences are also shown in the third and fifth cardiac cycles of FIG. 4. ‘Blank’ cardiac cycles containing no IR pulses or IR-DA sequences (e.g., the second and fourth RR cycles in FIG. 4) are typically provided as a delay between successive IR-pulse cycles to allow for further recovery of the magnetization before additional image data is obtained.
In this conventional cardiac triggered imaging sequence, the IR pulse is generated at a fixed interval following an R-wave, and the IR-DA sequence also occurs at a fixed interval following the R-wave when obtaining data for a single image over multiple cardiac cycles. The fraction of an interval between successive R-waves is commonly referred to as a cardiac ‘phase’; e.g., the midpoint between two R-waves can define a particular cardiac phase, the point in time that is ⅓ of the way between an R-wave and the next R-wave can represent another phase, etc. Such phases are substantially independent of the actual duration of the R-R interval (e.g., independent of the specific heartbeat rate), and can represent a particular shape configuration of the heart as it cycles through the repeated beating process.
In cardiac imaging, image datasets are typically obtained for a particular phase over a plurality of cardiac cycles so that the heart, which is rhythmically moving during the cardiac cycles, will have the same “shape” during each IR-DA sequence. Without such timing, different IR-DA sequences that are used to reconstruct a single image may obtain image data when the heart has different shapes within the cyclical heartbeat sequences, leading to corrupted image data. Similarly, pulsatile blood motion may lead to different locations and/or shapes of extra-cardiac tissues throughout the cardiac cycle. Thus, appropriate cardiac timing may be desirable to accurately image extra-cardiac tissues, such as the vasculature and other organs, which are affected by the blood motion or by the pulsing heart itself. Without such timing and triggered data acquisition, image data may be corrupted due to cardiac-induced motion of the imaged tissues.
Normal breathing by the subject or other bodily movement while MR image data is being collected can also lead to image corruption, because the region being imaged can move relative to the MRI apparatus. One technique to reduce such unwanted motion is a simple ‘breath-hold’ technique where the subject is instructed to hold their breath during the imaging procedure, to reduce or eliminate motion of the diaphragm and chest cavity. If the subject's breath can be held for several heartbeats, image data can be obtained that is not affected by motion of the diaphragm. However, many subjects do not hold their breath perfectly and a small non-negligible diaphragmatic drift may occur, leading to motion-based image corruption. Additionally, involuntary swallowing during a breath-hold can corrupt the images.
Another well-known technique for reducing the effects of subject motion during MR imaging involves the use of ‘navigator’ images, often referred to as ‘navigators.’ These navigators are typically single k-space lines acquired using “navigator echoes” or low-resolution images of a small portion of (or adjacent to) the region being imaged. Such navigators can be obtained at several points during the overall imaging sequence in a relatively short amount of time. Alignment of these navigator lines or images can be used to align the image data obtained temporally proximal to the navigators, which can provide a degree of correction when processing the image data for undesirable motion that occurred during image data acquisition.
The phase-sensitive inversion recovery technique (PSIR) was developed to overcome the non-linear relationship between magnetization and image brightness. PSIR is a common MR imaging pulse sequence and reconstruction technique that provides IR images with good T1 contrast, even if the inversion time TI between an IR pulse and the IR-DA data readout sequence has not been set optimally. This is an advantage over the standard IR sequence, where an incorrect TI value can result in poor image contrast or incorrect interpretations of clinical images. PSIR images may be ECG-triggered for cardiac imaging and for imaging of extra-cardiac tissues that are prone to have significant motion during the cardiac cycle due to pulsatile blood flow.
In PSIR, image pixel (or voxel) intensity is displayed on a grayscale where the maximum magnetization in the image dataset (typically infarcted myocardium when imaging the heart) is depicted as white in the acquired image, and the minimum magnetization in the image dataset (typically normal myocardium or a fluid with intrinsically long T1 value) is displayed as black. With PSIR image reconstruction, the phase information (+/− magnetization) is restored to the image by comparing changes in the phase of the image magnetization between paired (i.e., corresponding) image datasets. In practice, two paired datasets are acquired for each PSIR pulse sequence, the first of which is a conventional IR dataset (IR-DA), and the second being a reference phase MR dataset (REF). When the REF dataset is acquired, there is little or no perturbation to the longitudinal magnetization M0 of the imaged tissue resulting from a prior RF pulse (e.g., the tissue has substantially no magnetization preparation).
A schematic illustration of a conventional triggered PSIR data acquisition sequence is shown in FIG. 5. This is similar to the triggered IR sequence illustrated in FIG. 4, except that each IR pulse and subsequent IR-DA image data collection sequence is followed by a reference data collection sequence (REF) during the subsequent cardiac cycle.
An important aspect of the PSIR image reconstruction is spatial registration between the two datasets (IR-DA and REF), as corresponding pixels from both datasets are compared in order to restore the phase information. To minimize artifacts due to spatial misregistration, the two datasets are typically acquired in an identical manner. For segmented k-space acquisition, the identical portions of k-space are typically acquired during two consecutive heart beats at the same temporal window within the cardiac cycle (e.g., centered over the same cardiac phase as described earlier) for the IR-DA and REF datasets, and with identical spatial and temporal resolution. Thus, for cardiac imaging, the REF and IR-DA image datasets will correspond to the heart having the same shape as it cycles through periodic heartbeats. For extra-cardiac imaging, the REF and IR-DA image datasets will correspond to images of extra-cardiac tissues that are deformed in the same transient manner by pulsatile blood flow. As can be seen in FIG. 5, in a standard embodiment of PSIR, the time of two heartbeats is required to acquire the corresponding portions of both datasets.
FIG. 6 shows an exemplary PSIR image of the same cardiac location of the same patient as the IR image of FIG. 3. In this image, the infarct appears as bright white, with the normal myocardium being dark grey, such that the infarct can be clearly identified. Within a large range of TI values, the PSIR image will exhibit similar image contrast for cardiac tissues, because such contrast is based on the actual differences in magnetization for the various tissues, and not just on the differences in their magnitudes (or absolute values). In essence, PSIR imaging techniques provide a sign-corrected image that accounts for the polarity of the magnetization. Accordingly, the contrast in a PSIR image is much less sensitive to the selection of a particular TI value.
The PSIR magnetic resonance imaging technique can be implemented with both segmented and single shot readout schemes. In a segmented acquisition, the data acquisition for a single image is acquired as a plurality of portions (segments) of the image dataset that are distributed over multiple heart beats. The image data for each segment covers only a portion of k-space. The cardiac image shown in FIG. 6 is an example of an image reconstructed from a segmented PSIR acquisition. Each data segment is acquired as a paired segment, usually spanning two heart beats, with the latter part of each paired segment being a reference dataset (REF) used for phase identification as illustrated in FIG. 5. In a single shot acquisition, all IR image data are collected in one acquisition corresponding to a single readout train. In a subsequent heartbeat, all the reference data are acquired. As with the segmented implementation, both acquisitions (IR-DA and REF) are obtained during (e.g., the data acquisition sequence is centered at) the same cardiac phase (e.g., at the same time interval following the R-wave), but during consecutive heartbeats or cardiac cycles (as illustrated, e.g., in FIG. 5).
Conventional PSIR techniques, e.g., as illustrated in FIG. 5, are susceptible to imaging artifacts arising from spatial misregistration of the conventional data and reference phase data. Nonetheless, it is commonly assumed that by acquiring both datasets (IR-DA and REF) in an identical manner—e.g., in the same time point window within the cardiac cycle (centered over the same cardiac phase of two consecutive heart beats) and with the same spatial and temporal resolution—image artifacts are reduced to their lowest level, and that image artifacts would be worse if both datasets were not acquired in an identical manner. These common assumptions can lead to suboptimal clinical results in patients because of the extended time periods needed to obtain the PSIR paired datasets (especially when obtaining a plurality of such paired datasets in segmented acquisitions), which may introduce various motion artifacts during the extended imaging procedure.
Another limitation of conventional PSIR techniques is that they are not well-suited for single-shot imaging during free-breathing acquisitions. Some subjects have difficulty holding their breath and/or they may exhibit diaphragmatic drift; single-shot sequences can be helpful in reducing motion artifacts for such subjects as compared with segmented k-space acquisitions because of the shorter overall acquisition time needed. However, the requirement of acquiring two paired image datasets (IR-DA and REF) during PSIR imaging increases the likelihood of substantial breathing motion between the two datasets. Note that in single shot PSIR imaging, the IR-DA and corresponding REF image datasets are typically acquired two heartbeats (e.g. about 1.5 to 2 seconds) apart to allow for magnetization recovery after the magnetic saturation created by the long single-shot readout before acquiring the REF dataset. As a consequence, the final PSIR image can have motion artifacts due to spatial misregistration between IR-DA and REF datasets even though each of the two paired datasets was acquired in a single shot. Although moving the IR-DA and REF data acquisitions closer together in time could result in fewer motion artifacts, it is commonly thought that such a shortened interval is not possible in conventional PSIR imaging procedures because the two datasets need to be acquired at the same cardiac phase in separate heartbeats.
As noted earlier, a perfect breath hold can result in improved spatial registration of the IR-DA (inversion recovery) datasets and corresponding (paired) reference datasets in PSIR imaging when each dataset is acquired in the same cardiac phase, even though such paired dataset acquisition requires two separate cardiac cycles (heartbeats) as shown, e.g., in FIG. 5. Imperfect breath holds or free breathing imaging procedures, however, can lead to substantial shifts in cardiac position due to respiration-induced motion, thereby degrading spatial registration between the IR dataset and the corresponding reference dataset, leading to artifacts in the final PSIR image. In clinical practice, imperfect breath holds (or an inability to hold the breath for an extended period) are common, resulting in significant artifacts for the resulting PSIR images. Even when a subject can perform a perfect sustained breath hold, there may be ectopic heartbeats (such as a premature ventricular contraction) during the PSIR acquisition, which can also lead to spatial misregistration errors in the reconstructed PSIR image.
An important limitation of conventional PSIR techniques is that they are not well-suited for 3D (or 2D) respiratory-navigated procedures. In such imaging procedures, a respiratory navigator triggers the acquisition of image data. During a free-breathing respiratory-navigated sequence, respiratory navigators monitor the position of the diaphragm, and the associated acquired image data is only retained when the navigators determine the diaphragm is in the correct location.
In one type of respiratory-navigated PSIR, a single navigator is used for acquiring both datasets—the IR-DA dataset and the associated subsequent REF dataset. This approach can produce significant artifacts, because the reference dataset is usually acquired more than one second after the navigator data is acquired. If any motion of the subject occurs (e.g. breathing) between acquisitions of the IR-DA dataset and the reference dataset, spatial misregistration artifacts will occur even if the single navigator indicates that the image data is “good” data.
A second version of respiratory-navigated PSIR imaging uses a separate navigator for each of the inversion recovery and reference datasets. In principle, the addition of a second navigator can better account for the possibility of subject motion between the IR-DA and REF datasets. However, the need for a second navigator results in a significant lengthening of the overall scan time needed for data acquisition. Specifically, the scan time may be more than double that needed for data acquisition of a comparable 3D (or 2D) IR image without phase sensitivity information.
For example, the acquisition of paired datasets (e.g., a portion of k-space for segmented acquisitions, or 100% of k-space for single-shot acquisitions) during a PSIR imaging procedure will usually take far longer than two heartbeats because two distinct respiratory navigators have to accept data for the IR and reference datasets, respectively, and it is unlikely that the diaphragm will be in the same location for two consecutive heart beats. Scan times can become so excessive that this version of respiratory navigated PSIR is rarely attempted in clinical practice. To overcome this limitation, some navigated PSIR techniques do not navigate the reference data, assuming that the errors in misregistration will not greatly affect image quality. However, as discussed above, this assumption is usually invalid, and because of poor resulting image quality, this type of respiratory-navigated PSIR (with non-navigated reference datasets) is also rarely attempted in clinical practice.
Other techniques have been proposed for reducing the effects of respiratory or cardiac motion on MR image quality. For example, the so-called motion correction or “MOCO” technique does not reduce motion-related misregistration during image data acquisition, but instead attempts to compensate for such motion after the image data is obtained. This is achieved by post-processing, e.g., by non-rigid co-registration of images acquired during different parts of the respiratory cycle, cardiac cycle, or both. This technique works only with 2D datasets, and cannot correct for through-plane cardiac position shifts that typically occur with free breathing or poor breath-holding.
Accordingly, it would be desirable to have a system and method for magnetic resonance imaging that addresses some of the shortcomings described above, and that may further provide improved PSIR imaging of tissue affected by cardiovascular pulsatile motion by reducing the presence of motion artifacts and other dataset misregistration effects.